1. Field of the Invention
The present invention relates to an optical tomography apparatus that irradiates a low coherence measuring light beam onto a measurement target to obtain tomographic images of the measurement target. Particularly, the present invention relates to an optical tomography apparatus that obtains images of the surface and the fine structures within the measurement target, based on a reflected light beam, which is the measuring light beam reflected by the measurement target.
2. Description of the Related Art
As conventional methods for obtaining tomographic images of measurement targets, such as living tissue, methods that obtain optical tomographic images by OCT (Optical Coherence Tomography) measurement have been proposed. As an example of an OCT measurement method, the TD-OCT (Time Domain Optical Coherence Tomography) measurement has been proposed (refer to Japanese Unexamined Patent Publication Nos. 6(1994)-165784 and 2003-139688). The TD-OCT measurement is a type of light interference measurement method that utilizes the fact that light interference is detected only when the optical path lengths of divided light beams, that is, a measurement light beam and a reference light beam, match within a range of coherence length of a light source. That is, in this method, a low coherent light beam emitted from a light source is divided into a measuring light beam and a reference light beam, the measuring light beam is irradiated onto a measurement target, and the measurement light beam reflected by the measurement target is led to a multiplexing means.
In the TD-OCT measurement, the measuring position (measuring depth) within the measurement target is changed, by changing the optical path length of either the reference light beam or the measuring light beam. Thereby, a one dimensional tomographic image in the direction of the optical axis is obtained. For example, the TD-OCT apparatus disclosed in Japanese Unexamined Patent Publication No. 6(1994)-165784 comprises an optical system that causes a reference light beam emitted from an optical fiber to be reflected by a mirror. The optical path length of the reference light beam is adjusted by moving the mirror in the direction of the optical axis of the reference light beam. In addition, the irradiation position of a measuring light beam, which is irradiated on a measurement target, is scanned in a direction perpendicular to the optical axis thereof, thereby enabling obtainment of two dimensional tomographic images based on two dimensional reflected optical intensities. Further, by scanning the irradiation position of the measuring light beam two dimensionally perpendicular to the optical axis thereof, three dimensional tomographic images can be obtained, based on three dimensional reflected optical intensities.
As another OCT measurement method, a method that obtains optical tomographic images by SD-OCT (Spectral Domain Optical Coherence Tomography) measurement has been proposed (refer to U.S. Pat. No. 6,377,349). In an SD-OCT apparatus, a wide band low coherence light beam is divided into a measuring light beam and a reference light beam. The optical path lengths of the measuring light beam and the reference light beam are substantially matched, then the two light beams are caused to interfere with each other, to form a coherent light beam. Thereafter, the coherent light beam is decomposed into different frequency components by a spectral decomposing means. An array type photodetector measures the intensity of each frequency component of the coherent light beam. The coherent spectral waveform obtained by the photodetector undergoes Fourier transform at a computer, to obtain one dimensional tomographic data in the direction of the optical axis, without physically changing the optical path length. By scanning the measuring light beam in directions perpendicular to the optical axis, two dimensional and three dimensional tomographic images can be obtained.
Further, the SS-OCT (Swept Source Optical Coherence Tomography) method has been proposed in U.S. Pat. No. 5,956,355. In the SS-OCT method, a coherent light beam, of which the frequency is temporally varied, is emitted instead of a low coherence light beam. The coherent light beam is detected, and reflection intensities at depth positions within a measurement target are calculated, based on interferograms of optical frequency regions. Then, tomographic images are generated employing the calculated reflection intensities.
These OCT apparatuses have been developed and are in use in the field of ophthalmology. Following the use of OCT apparatuses in the field of ophthalmology, research and development are underway for application in endoscopes. In the initial stages of development, the 0.8 μm band had been employed as the wavelength of the light sources of the OCT apparatuses (refer to W. Drexler et al., Optics Letters Vol. 24, No. 17, pp 1221-1223, 1999.). This wavelength band was selected as a result of considering absorption properties of living tissue. FIG. 1A is a graph that illustrates light absorption coefficients of water, blood, melanin, and epidermis. FIG. 1B is a graph that illustrates the absorption coefficients of water with respect to light having wavelengths between 0.7 μm and 1.6 μm. From the graph of FIG. 1B, it can be seen that the peak of absorption occurs at 0.98 μm and at 1.2 μm. In addition, the broken line in the graph of FIG. 2 is a graph that represents absorption loss in living tissue, based on the absorption coefficients. From the graph of FIG. 2, it can be seen that light within the 0.8 μm band has the smallest amount of absorption loss. For this reason, it was considered that light within the 0.8 μm band has the highest transmissivity with respect to living tissue, enables deeper measurement depths, and is most suited for OCT apparatuses.
However, it has been found recently that scattering properties also limit measurement depths in OCT apparatuses. This is because OCT apparatuses detect backscattered reflected light beams from within living tissue. Rayleigh scattering is common within living tissue. In Rayleigh scattering, the scattering intensity is inversely proportionate to wavelength to the fourth power. The dotted line in the graph of FIG. 2 represents scattering loss within living tissue. The total loss, represented by the solid line in the graph of FIG. 2, is the sum of the absorption loss and the scattering loss.
From the graph of FIG. 2, it can be seen that the wavelength band, at which total loss is minimal, is the 1.3 μm band. For this reason, after OCT apparatuses for ophthalmology were realized, research and development for OCT apparatuses to be applied to endoscopes, which require deeper imaging depths, are being performed with the 1.3 μm band as the wavelength of light sources therein (refer to Japanese Unexamined Patent Publication No. 2003-139688).
The purpose for applying an OCT apparatus to an endoscope is to enable definitive diagnoses within living organisms, and to diagnose the depth of tumor invasion of mucosal cancer (m cancer) and submucosal cancer (sm cancer). Hereinafter, the procedure of endoscopic diagnosis of cancer will be briefly described. First, a diseased portion is discovered within a normal observation image, and whether the disease is cancer or another illness is discriminated. This preliminary diagnosis is based on the experience of a physician, after which tissue from a portion estimated to be cancerous is collected and subjected to a biopsy, to obtain a definitive diagnosis. For this reason, it is presently difficult to obtain definitive diagnoses during examination with an endoscope. In the case that a diseased portion is definitively diagnosed as cancer, the depth of tumor invasion is diagnosed by endoscopic examination, in order to determine a treatment strategy. Commonly, cancers present themselves in the mucoepidermis, and metastasize in the horizontal direction and in the depth direction, as the disease progresses. As illustrated in FIG. 3, the structure of a stomach wall is constituted by: a membrana mucosa (m) layer; lamina muscularis mucosae (MM); a submucosal (sm) layer; tunica muscularis ventriculi; and a serous membrane. Cancers which are present only in the membrana mucosa layer are designated as m cancers, and cancers which have penetrated to the submucosal layer are designated as sm cancers. Treatment protocols differ between m cancers and sm cancers. Blood vessels and lymph systems are present in the submucosal layer, and there is a possibility of metastasis in the case of sm cancers. Therefore, surgical procedures are required. On the other hand, there is no possibility of metastasis in the case of m cancers. Therefore, m cancers are removed by endoscopic procedures. For this reason, it is necessary to discriminate whether cancers are m cancers or sm cancers. Specifically, it is important to be able to evaluate whether the layer structure of the lamina muscularis mucosae layer is maintained or destroyed, in an image. Presently, application of ultrasound imaging techniques is being considered, with the objective of diagnosing the depth of tumor invasion. However, the resolution of ultrasound imaging is only about 100 μm in the axial direction, which is insufficient to visualize the MM layer. In addition, in m cancers which have progressed, lymph follicles are formed under the MM layer, thereby causing the cancerous portions and the lymph follicles to be imaged integrally, and m cancers may be misdiagnosed as sm cancers. For this reason, an imaging method having a resolution of 10 μm or less in the axial direction is desired, to enable accurate diagnosis of the depth of tumor invasion.
Meanwhile, the resolutions of TD-OCT and SD-OCT apparatuses in the optical axis direction are determined by the coherence length of the light sources thereof. That is, it is not generally possible to obtain resolution less than the coherence length of the light source. For this reason, a light beam having a coherence length of 10 μm or less is necessary to obtain high resolution of 10 μm or less. The coherence length Δz of low coherence light is proportionate to the square of the central frequency and inversely proportionate to the spectrum width thereof. The coherence length Δz can be expressed by the following formula:Δz=(21n2/Π)·(λc2/Δλ)wherein                λc: central wavelength        Δλ: spectrum width        
For this reason, it is necessary to broaden the spectrum width Δλ in order to decrease the coherence length. Meanwhile, it was found that the influence of dispersion needed to be considered, if the spectrum width Δλ was broadened (refer to Y. Wang et al., Optics Express Vol. 11, No. 12, 2003, pp 1411-1417, 2003.).
In a Michaelson interferometer, as a light beam propagates through a sample, phase shift occurs, and a coherent signal waveform changes as a result. If the coherent signal waveform is designated as φ(w) and the spectrum waveform of the light source is a Gaussian distribution, autocorrelation functions can be expressed as:
                              δ          t                =                              δ                          t              ⁢                                                          ⁢              0                                ·                                    {                              1                +                                                      〈                                                                                            ⅆ                          2                                                ⁢                                                  φ                          ⁡                                                      (                            w                            )                                                                                                                      ⅆ                                                  w                          2                                                                                      〉                                    ⁢                  δ                  ⁢                                                                          ⁢                                      w                    4                                                              }                                      1              2                                                          (        1        )                                K        =                              δ            t                    /                      δ                          t              ⁢                                                          ⁢              0                                                          (        2        )                                D        =                                            -                              w                0                2                                                    2              ⁢              π              ⁢                                                          ⁢              c                                ·                                                    ⅆ                2                            ⁢                              φ                ⁡                                  (                  w                  )                                                                    ⅆ                              w                2                                                                        (        3        )            wherein                δt: 1/e1/2 width of the autocorrelation function        δt0: 1/e1/2 width of the autocorrelation function when D=0        δw: 1/e1/2 width of the optical spectrum        w0: central frequency of the optical spectrum        K: broadening ratio due to the influence of dispersion        
FIG. 4 is a graph that illustrates calculated results (represented by the solid line) of formula (3) above and actual measured values (represented by the triangles). Dispersion D is zero when the wavelength of the light beam is 1.0 μm. It can be seen from the graph of FIG. 4 that the influence of dispersion becomes greater as the wavelength becomes greater than or less than 1.0 μm.
FIG. 5 is a graph that illustrates measured values and simulation results of the relationship between the distance of propagation (depth of water) and broadening ratios, when low coherence light beams having wavelengths of 1.32 μm (spectrum width: 76 nm) and 0.94 μm (spectrum width: 75 nm) propagate through water.
In the aforementioned document, Y. Wang et al. conclude that it is preferable to employ low coherence light having a central wavelength of 1.0 μm in OCT apparatuses, in the case that the coherence length of the low coherence light beam is short.
The resolution in the optical axis direction is defined by the wavelength sweep width Δλ of the coherent light beam emitted by the light source in SS-OCT apparatuses as well. For this reason, the wavelength sweep width Δλ needs to be widened, in order to increase the resolution in the optical axis direction. However, if the wavelength sweep width Δλ is widened, it becomes necessary to consider the effects of scattering, as described above.
However, when an OCT apparatus that employs low coherence light of coherent light, of which the frequency is temporally varied, is used to obtain an optical tomographic image of an organism, there are cases in which the wavelength band of the measuring light beam includes wavelengths which are readily absorbed by living tissue. In these cases, the spectral shape of the reflected light beam changes due to the light absorption by the living tissue, and side bands or side lobes appear in the autocorrelation function, generating pseudo signals that reduce the S/N ratio of the optical tomographic image. As illustrated in FIG. 1B, peaks in the absorption coefficient of water, which is the main constituent of living tissue, occur at wavelengths of 0.98 μm and 1.2 μm.
FIGS. 6A and 6B are graphs that represent simulations of a light beam having a central wavelength of 1.0 μm, a wavelength band width of 100 nm, and a Gaussian distribution propagating through water. FIG. 6A represents the changes in spectrum shape, and FIG. 6B represents Fourier transform waveforms of each spectral waveform. Note that the solid lines represent waveforms which are not transmitted through water; the long/short dashed lines represent waveforms which have been transmitted through 2 mm of water; the long/short/short dashed lines represent waveforms which have been transmitted through 4 mm of water; and the broken lines represent waveforms which have been transmitted through 8 mm of water. It can be seen from the graphs of FIGS. 6A and 6B that in the case that a measuring light beam having a central wavelength of 1 μm and a wavelength band width of 100 nm propagates through water, influence of absorption at 0.98 μm greatly changes the spectrum shape. As a result, side bands appear in the autocorrelation waveform, pseudo signals are generated, and the quality of the optical tomographic image deteriorates.
In the aforementioned document by Y. Wang et al., it is disclosed that influence due to scattering is observed when optical tomographic images are obtained employing low coherence light having a coherence length of approximately 10 μm (λc2/Δλ=23). In addition, Y. Wang et al. disclose that it is preferable to set the central wavelength of low coherence light in the vicinity of 1.0 μm in cases that influence due to dispersion is observed. However, there is no disclosure regarding a central wavelength λc nor a wavelength band width Δλ that avoids influence due to absorption at the 0.98 μm and 1.2 μm wavelengths.